1. Field of the Invention
This invention relates generally to an apparatus and method for medical examination, in particular, a lutetium oxyorthosilicate (LSO) or light-output equivalent positron emitting tomography (PET) detector.
2. Description of Related Art
The American Cancer Society has predicted that there will be more than 181,000 new breast cancer cases and more than 40,000 deaths from breast cancer in the United States in 2000. [American Cancer Society, xe2x80x9cCancer Facts and Figuresxe2x80x941999,xe2x80x9d American Cancer Society, Atlanta, Ga. (1999).] Breast cancer is also the second leading cause of cancer death in women. Currently, mammography and physical breast examination, provide the two most effective methods for screening potential breast cancer patients. Although mammography allows the detection of very small, non-palpable lesions, it has a limited diagnostic accuracy for detecting cancer and image interpretation is subject to considerable inter-observer and intra-observer variability. The incidence of positive biopsies performed after mammographic findings ranges from 9% to 65%, with most investigators reporting a 15 to 30% positive biopsy rate. The sensitivity of detection by mammography drops considerably in women with dense, fibrocystic breasts.
Microcalcifications, one of the classic signs of occult malignancies, have a low predictive value of only 11.5% for the presence of cancer. The predictive value of masses that are thought to definitely represent malignancies is about 74%, but masses thought to be possibly malignant turn out to be carcinoma in only 5.4% of the cases. [M. Moskovitz, xe2x80x9cThe predictive value of certain mammographic signs in screening for breast cancer,xe2x80x9d Cancer, 51, 1007-1011 (1983)]. Also, several studies have reported substantial variability among radiologists in interpretation of mammographic examinations. [K. Kerlikowske, et al., xe2x80x9cVariability and accuracy in mammographic interpretation using the American College of Radiology Breast Imaging Reporting and Data System,xe2x80x9d J. Natl. Can. inst., 90, 1801-1809 (1998)]. Therefore, mammography is a useful screening tool for detecting cancer, but it is limited by a large number of false positive tests, which result in unnecessary biopsies. Mammography is also limited by a considerable number of false negative tests, which result in the missed diagnosis of cancer.
It is also possible to use radio-pharmaceutical and radio-nuclide imaging to detect cancers, such as [18F]fluoro-2-deoxy-D-glucose (FDG). FDG is a radioactive analog of glucose, which is phosphorylated and trapped within cells. After a patient receives a dose of FDG, she may be examined with a detector that senses the gamma rays produced by 18F. Positron emission tomography (PET), using FDG as a tracer of tumor glucose metabolic activity, is an accurate non-invasive imaging technology which probes tissue and organ function rather than structure. [See U.S. Pat. No. 5,453,623 and U.S. Pat. No. 5,961,457]. The increased rate of glycolysis in neoplastic cells, independent of the oxygen concentration present, has been previously reported. [O. Warburg, xe2x80x9cOn the origins of cancer cells,xe2x80x9d Science, Vol. 123, 309-314 (1956) and U.S. Pat. No. 5,969,358]. This information is fundamental to the utility of FDG for imaging human neoplasms.
Whole body PET scanners are used clinically to diagnose and to stage a wide variety of cancers. [C. K. Hoh, et al., xe2x80x9cPET in oncology: will it replace the other modalities?xe2x80x9d Sem. Nucl. Med., 27, 94-106 (1997)]. PET scanners detect breast cancer with sensitivities between 70 and 90% and with specificities of 84-97%. [N. Y. Tse, et al., xe2x80x9cThe application of Positron Emission Tomographic imaging with fluorodeoxyglucose to the evaluation of breast disease,xe2x80x9d Ann Surg., 216, 27-34 (1992); O. E. Nieweg, et al., xe2x80x9cPositron Emission Tomography of Glucose Metabolism in Breast Cancer: Potential for Tumor Detection, Staging, and Evaluation of Chemotherapy,xe2x80x9d Ann. N. Y. A. Sci., 698, 423-448 (1993); and R. L. Wahl, et al., xe2x80x9cPrimary and Metastatic Breast Carcinoma: Initial Clinical Evaluation with PET with the Radiolabeled Glucose Analogue 2-[F-18]-Fluoro-2-deoxy-D-glucose,xe2x80x9d Radiology, 179, 765-770 (1991)]. A high diagnostic accuracy of PET imaging for staging of axillary lymph node involvement has also been reported. [L. Adler, et al., xe2x80x9cAxillary lymph node metastases: screening with F-18 2-deoxy-2fluoro-D-glucose (FDG) PET,xe2x80x9d Radiology, 203, 323-327 (1997)]. The. lower than desired diagnostic accuracy reported for PET imaging is due to relatively poor accuracy for detecting tumors of less than 1 cm in size. [N. Avril, et al., xe2x80x9cMetabolic characterization of breast tumors with positron emission tomography using F-18 fluorodeoxyglucose,xe2x80x9d J Clin. Onc., 14, 1848-1857 (1996)].
Most PET imaging technology is currently based on scintillation detectors. Radiation detection begins by injecting isotopes with short half-lives into a patient""s body. The isotopes are absorbed by target areas within the body, causing the isotope to emit positrons that are detected when they generate gamma rays. When in the human body, the positrons collide with electrons and the two annihilate each other, releasing gamma rays. The emitted rays move in opposite directions, leave the body and strke the array of radiation detectors. In the majority of commercial PET systems, a xe2x80x9cblockxe2x80x9d design composed of a high-density, partially-segmented (for weighted light sharing) scintillation crystal (bismuth germanate) is coupled to four photomultiplier tubes (PMTs). [M E. Casey, et al., xe2x80x9cA multicrystal two dimensional BGO detector system for positron emission tomography,xe2x80x9d IEEE Trans. Nucl. Sci., 33, 460-463 (1986) and S. R. Cherry, et al., xe2x80x9cA Comparison of PET Detector Modules Employing Rectangular and Round Photomultiplier Tubes,xe2x80x9d IEEE Trans. Nucl. Sci., 42, 1064-1068 (1995) and U.S. Pat. No. 5,453,623]. In this design, the scintillation crystal is subdivided into semi-discrete crystals by incomplete cuts which are filled with reflecting material. The PMTs are not position-sensitive and rely on the different depths of the cuts in the scintillation crystal to yield a light distribution on the PMT""s which varies linearly with interaction position across the detector. A problem with the block design of current PET systems is that the intrinsic spatial resolution and the spatial sampling of the block is determined by the size of the individual crystals. In order to improve the intrinsic spatial resolution the size of the crystals needs to be reduced. However, with the block design it becomes difficult to decode smaller crystals. Another problem inherent to the block design PET system is that it is fairly bulky, because of the large dimensions of most single channel PMTs.
More recently, high resolution, high sensitivity PET detectors have been constructed by directly coupling the scintillator material 4 to a compact, low-cost, position-sensitive PMT (PS-PMT). By coupling small discrete scintillator elements 4 directly onto the active area of the PS-PMT, one maximizes light transmission from the scintillator 4 to the PS-PMT. [J. J. Vaquero, et al., xe2x80x9cPerformance Characteristics of a Compact Position-Sensitive LSO Detector Module,xe2x80x9d IEEE Trans. Nucl. Sci.,17, 967-978 (1998) and R. Pani, et al., xe2x80x9cMulti-PSPMT scintillation camera,xe2x80x9d IEEE Trans. Nucl. Sci., 46, 702708(1998) and U.S. Pat. No. 5,864,141]. However, these PS-PMT""s have a significant inactive area at the edges. Using the direct coupling method and tiling many detectors together to form planar arrays, therefore, produces large gaps between the detector modules 20 because the effective or active area 10 of the PMT 8 does not span the full physical dimensions of the face of the tube (FIG. 1a). This reduces system sensitivity and sampling and causes problems in the reconstruction of the data. Therefore, it is desirable to develop some sort of tapered light guide 12 to eliminate these gaps and to form large continuous arrays (FIG. 1b). [R. Pani, et al., supra.]
A PET camera based on discrete LSO scintillator elements and a fixed ring geometry has been reported. [W. Moses, et al., xe2x80x9cPET camera designs for imaging breast cancer and axillary node involvement,xe2x80x9d J. Nucl. Med., 36, 69P (1995) and U.S. Pat. No. 6,040,580]. However, the flexibility of the planar detector arrays with variable separation of the present invention offers advantages in the clinical setting over PET systems in a fixed ring geometry.
Conventional PET imaging devices are designed to image cross sections of the entire body. Although functional imaging with PET is a promising technique in conjunction with x-ray mammography for breast cancer patient management, there are several disadvantages to employing a whole body PET scanner for the detection of malignant breast tumors. The first disadvantage is that the whole body PET system is limited by the spatial resolution and sensitivity. [N. Avril, et al., supra]. Whole body PET systems typically yield reconstructed images with a resolution of 8-15 mm, depending on the injected dose, imaging time, and intrinsic resolution of the scanner. The effect of this resolution limit is that radioactivity is underestimated.
The second disadvantage of a conventional whole body PET is the high cost of the examination. Whole body PET is an expensive technology, and is generally only available in the larger medical facilities in the United States.
A third disadvantage of a conventional whole body PET scanner is that the PET scanner provides metabolic images of breast cancer patients with several shortcomings related to the general-purpose nature of these systems, e.g., in whole body scanners the detectors are typically 20-30 cm away from the breast or axilla, which reduces sensitivity. Conventional scanners also have relatively large detector elements (greater than 4 mm), which limits spatial resolution.
The present invention is an apparatus and method for examining a body part. In particular, the present invention is directed to a dedicated PET system for breast imaging or imaging other body parts, such as the head, neck, liver, heart, lungs and other extremities, which overcomes the limitations of prior detectors and improves the overall diagnostic quality of the images.
The positron emission tomography imaging apparatus of the present invention is a dedicated mammary and axillary region PET imaging system and comprises at least two large-area planar scintillation detector plates composed of 25, a 5xc3x975 array, of modular detectors. The detectors include an array of scintillation crystals, a plurality of photomultiplier tubes positioned adjacent the plurality of scintillation crystals, and a lightguide having an end positioned adjacent to the array of scintillation crystals and having an opposing end adjacent to the photomultiplier tubes.
The planar scintillation detector plates 22 operate in coincidence and have about a 15xc3x9715 cm2 surface area, giving complete coverage of the breast in a single view. The detector can be mounted on a flexible gantry, allowing the inter-detector separation to be varied from about 10 cm up to about 50 cm, and also allowing the detectors to rotate to collect tomographic information.
A method for examining a body part is also described and comprises providing an internal image of the body part including, a positron emitting radioisotope and a positron recording apparatus between which the body part is to be disposed; and placing at least two detector plates 22, each plate comprised of at least one detector, said detector having a scintillator coupled to one end of a lightguide, the opposing end of said lightguide coupled to a photomultiplier tube, said detector is capable of detecting gamma-rays emitted by the radioisotope infiltrated into the body part in an adjacent relationship with said recording apparatus for providing the internal image.
The system of the present invention allows for adjustable detector separation to accommodate all patients and permits imaging of the axillary region. The adjustable detector plate separation also allows closeness to the area being imaged, thereby increasing the system sensitivity. The detector plates 22 or arrays are large enough to scan an entire breast in one imaging setup. The flexible scanner geometry allows planar, limited angle, filtered back projection or iterative image reconstruction techniques to be implemented.
The present invention has a number of important advantages over conventional PET scanners. The present invention brings the detectors in close to the breast or axilla, resulting in a large increase in the system sensitivity (the fraction of emitted gamma ray pairs that are detected) which is due to the increase in solid angle. This increase in sensitivity allows for an improved image signal-to-noise and/or image resolution. The increase in sensitivity also results in a reduced imaging time (increasing patient throughput) and/or allows for a smaller injected dose of FDG. Furthermore, the present invention allows for the gamma ray pairs to only pass through the breast to be detected, and does not require them to pass through the entire cross-section of the chest. Therefore, tissue attenuation is reduced and the correction for gamma-ray attenuation can be based on simple geometric calculations. Lastly, because relatively few detector modules 20 are needed to construct the scanner of the present invention (about 50 in the proposed system versus 250-300 in a whole-body PET scanner), the overall cost of the technology is dramatically reduced. This, in conjunction with the widespread availability of FDG from the growing network of PET radio-pharmaceutical distribution centers, results in PET becoming a viable diagnostic tool for breast cancer patient management.
This system offers several advantages when compared with existing dedicated PET systems for breast imaging and conventional whole-body PET scanners. Firstly, LSO scintillators provide important advantages over dedicated systems that use BGO crystals. Lutetium oxyorthosilicate (LSO) scintillators have a decay time of 40 ns which provides count-rate performance advantages. Thus, LSO has similar stopping power to BGO, but produces five times as much scintillation light and has a seven-fold shorter decay time, which enables the detector of the present invention to operate successfully in the high singles count rate environment expected in breast imaging, due to nearby activity from the heart and liver. The increased light output allows good timing and energy resolution improving image quality by reducing the influence of randoms and scatter.
Secondly, the use of an optical fiber taper allows detector modules 20 to be tiled together in planar arrays (with no gaps) which produce detector plates 22 of any desired size. For example, the 15xc3x9715 cm2 detector plates 22 of the present invention provide a large field of view which provide for better coverage and a shorter imaging time. Furthermore, these plates 22 are sensitive and maintain their resolution right to the very edges, allowing the closest possible imaging of the chest wall and imaging of the entire body part or breast. [C. J. Thompson, et al., xe2x80x9cPositron emission mammography (PEM): A promising technique for detecting breast cancer,xe2x80x9d IEEE Trans. Nucl. Sci., 42, 1012-1017 (1995); C. J. Thompson, et al., xe2x80x9cFeasibility study for positron emission mammography,xe2x80x9d Med. Phys., 21, 529-537 (1994); J. L. Robar, et al., xe2x80x9cConstruction and calibration of detectors for high resolution metabolic breast imaging,xe2x80x9d Nucl. Instrum. Methods Phys. Res. A, 392, 402-406 (1997); I. Weinberg, et al., xe2x80x9cPreliminary results for positron emission mammography: Real-time functional breast imaging in a conventional mammography gantry,xe2x80x9d Eur. J. Nucl. Med., 23, 804-806 (1996); R. Freifelder, et al., xe2x80x9cDedicated PET scanners for breast imaging,xe2x80x9d Phys. Med. Biol., 42, 2463-2480 (1997); and G. Hutchins, et al., xe2x80x9cEvaluation of prototype geometries for breast imaging with PET radiopharmaceutical,xe2x80x9d J. Nucl. Med., 36, 69P (1995)].
The present invention is also surprisingly inexpensive, as the tapered optical fiber bundles used in this work are a fraction of the cost of the very high resolution tapers used in conventional CCD based imaging systems. Lastly, the large-area plate geometry with variable separation allows unprecedented flexibility for clinical applications. The detector separation can be adjusted to suit the patient geometry and planar, limited angle or full tomographic datasets of the breast can be acquired. Planar images of the axilla may also be acquired with our proposed system. Bringing the detectors in close to the object of interest will improve sensitivity relative to whole-body PET scanners, and the resolution and timing performance of our detector modules 20 has been demonstrated to be superior to that measured in whole-body PET detectors. The goal of the maxPET system is to aid in breast cancer patient management by assisting in imaging patients with dense, fibro-glandular breasts, detecting axillary lymph node metastases without surgery and monitoring chemotherapy effectiveness.
These and other features, aspects, and advantages of the present invention will become better understood with regard to the following detailed description, appended claims, and accompanying drawings.